The field of coronary angioplasty and stenting has made dramatic progress in treatment of coronary heart disease through at least three generations of product technology. However, each generational advancement has been accompanied by a new challenge or failure mode of the therapy. Balloon angioplasty therapy improved acute luminal flow, but vessel recoil and remodeling resulted in high restenosis rates. Bare metal stenting lowered restenosis rates and minimized abrupt closure events, but restenosis rates were still high due to stent mechanical injury and resulting smooth muscle cell (SMC) migration and proliferation into the lumen. Drug eluting stents cut the retreatment rate again significantly by addressing the SMC proliferation with a pharmaceutical agent, but again was accompanied by a “new” complication, late stent thrombosis (LST) and the accompanying extended use of anti-coagulants. LST is associated with high mortality rates, although the frequency of the events is relatively low. The apparent factors driving this serious complication appear to be the loss of vaso-motion and delayed healing of a functional endothelium.
Attempts to use magnesium and its alloys as a temporary implant biomaterial, including in cardiovascular stents, have been hindered by poor control over the rate and uniformity of the metal's degradation (metallic corrosion rate), fragmentation and absorption processes in local tissue. Previous attempts to control degradation or corrosion rates have focused on alloying with rare earth and other heavy metal elements of unknown biocompatibility, yielding slower metallic corrosion rates but unproven benefits in clinical performance. Although these approaches have merit for non-medical applications such as commercial or aerospace castings, they are sub-optimal for an absorbable implant grade material that will eventually be fully metabolized by the host tissue, releasing alloying elements of unknown biocompatibility. Furthermore, conventional approaches to corrosion control of magnesium alloys have focused solely on preventing the initial mechanical failure of the given article by retarding the degradation process either by a surface passivation layer, or changing the local corrosion potential of the alloy. No consideration has been given to controlling the process of fragmentation, disintegration and absorption following the initial mechanical failure. For many implant applications, the timing and nature of the full degradation process, starting with the as-implanted metal article to the final clearance of the alloy mass and its degradants from the anatomical site, is critical regarding the performance of the medical device.
For absorbable metal implants, the corrosion process and ultimate mechanical properties are strongly dominated by the polycrystalline grain structure of the metal. Corrosion can proceed along grain boundaries due to localized galvanic reactions between Mg and more noble metals that are excluded from the crystal lattice during solidification from the melt. Cavitation and cracking can start at the grain boundaries due to the cyclic fatigue from pulsatile loading in the artery, resulting in gross mechanical failure of the implant long before a significant volume fraction of Mg has experienced corrosion. This can significantly shorten the implant's functional life, i.e., the period of time where the implant is mechanically intact and load bearing, and extend the time when large metallic fragments can cause injury and inflammation at the implant site.
One such implant application is absorbable metal stents for vascular or luminal scaffolding, such as stents for treatment of coronary artery disease. In this application, the stents provide temporary scaffolding through the healing process related to the local injury caused by the high pressure angioplasty balloon used to open the stenosed or partially blocked artery. The metal scaffold is typically required only for a period of days to weeks to prevent abrupt closure of the vessel from spasm, minimize elastic recoil, and as a substrate to deliver a controlled release drug-polymer formulation to the site of injury. After this period, any remnant of the alloy or its degradants may be a liability, since it can act as a foreign body prolonging an inflammatory response and delaying healing. Furthermore, if the stent remnants remain present in the lumen in solid form through the period of extracellular matrix deposition and scar formation, then the stent remnants themselves become a source of lumen obstruction and participate in a new form of restenosis unknown to conventional permanent stents.
An alternative design approach towards absorbable stents utilizes highly crystalline absorbable polymers such as PLLA for the structural elements of the stent scaffold. This approach has a more controlled degradation process, but suffers from low radial stiffness that is needed to open the artery, i.e., so called acute gain, and limited ductility making stent-artery sizing problematic.
The current standard of care for treating most de novo coronary lesions is the implantation of a permanent implant known as a drug eluting stent or DES. The DES is a third generation angioplasty device for treating coronary stenosis, with significantly lower re-intervention rates than either bare metal stents or balloon angioplasty. This generation technology is a permanent implant, typically comprising a high strength and high radiopacity metal such as cobalt chrome or platinum enriched stainless steel, coated with a formulation of an anti-proliferative drug in a controlled release polymer.
The next generation of technology is a fully absorbable DES, i.e. the entire mechanical scaffolding (stent) and the drug formulation is broken down in the body and absorbed. The working hypothesis is that any permanent foreign body at the site can prolong inflammation and delay healing and restoration to its native state. The major complication associated with drug eluting stents is late stent thrombosis, which is believed to result from this delayed healing.
The primary focus of fully absorbable stents has been on achieving the necessary hoop strength and stiffness to bear the high mechanical stresses in the coronary arteries, but a second key characteristic that is required is radio-opacity to enable the physician to visualize the stent after implantation. Since the two primary materials used in experimental absorbable stents, L-poly lactic acid (pLLA) and magnesium alloys, are both essentially radio-transparent, small disc-shaped radio-markers comprised of platinum, platinum-iridium, or tantalum are typically integrated into the end of the laser cut stent body. If there are 2 or 3 radio-markers on each end of the stent, then the location and level of deployment can be visualized by angiography during a procedure, even if the bulk of the stent is not radio-opaque. This is a well-established approach for nickel-titanium permanent stents which possessed low intrinsic radio-opacity.
The problems with this conventional approach for metallic absorbable stents, such as magnesium-based alloys, are that fragmentation occurs within weeks, and the relatively large radio-markers (approximately 1.0 mm diameter by 125 micron thick) may migrate from the implantation site and become emboli, potentially resulting in a serious infarction of distal coronary vessels. Ideally, the radio-marker for an absorbable magnesium stent would be self-sufficient regarding prevention of migration, and its safety not be compromised through magnesium fragmentation process.